The present invention relates generally to helmets, particularly helmets used to protect the head of a user participating in sports, such as football, or other activities. More particularly, the present invention comprises an improved helmet system for protecting a user from sustaining concussions and other head injuries.
A key function of sports helmets and football helmets in particular, is to reduce the occurrence of brain concussions. Concussion is the term used for mild traumatic brain injuries, MTBIs for short. Despite the “mild” descriptor, concussions are serious injuries and their effect if more than one is experienced by a player become cumulative and may lead to chronic traumatic encephalopathy, or CTE, with reduced brain function in later life. Plus recent evidence indicates that those with CTE may be fifty times more likely to get amyotrophic lateral sclerosis, or ALS, than the average population (Scientific American, February 2012). The problem today has become nearly epidemic—with an estimated 300,000 football concussions a year among youth, high school, college, and NFL players. Moreover, due to players concealing their injuries and coaches and trainers failing to detect them, many experts believe that number could be low by a factor of two. To counter the concussion problem, the NFL, the colleges, and the helmet manufacturers have attempted some or all of the following: improving the helmet designs; enforcing harsh penalties and severe fines for spearing or other intentional helmet to helmet contacts; identifying concussed players and keeping them sidelined long enough for symptoms to fully subside (sometimes several weeks); trying to better quantify the peak linear and angular acceleration levels of the skull that can lead to concussions; and in a combination of the latter two, measuring the accelerations in real time utilizing multiple miniature accelerometers located against the skull inside the helmets, with the skull acceleration waveforms being transmitted in real time to the sidelines so any player receiving a potential concussion level impact can be immediately identified and removed from the game to be administered predetermined concussion symptom checks, a test which the player must pass before being allowed to reenter the fray. A significant effort has also been made to come up with an optimum metric for characterizing skull impact levels that would accurately predict a resulting concussion. This task began several years ago with the severity index, SI; then the head impact criteria HIC; then head impact power HIP; and most recently the brain impact criteria, BIC and others. However, none of these metrics has yet been shown to be significantly more successful at predicting a concussion than the combination of the maximum linear acceleration value and the maximum angular acceleration value, where the current NFL threshold value being used for the former is 79 Gs, and the current NFL threshold value being used for the latter is 5,757 radians/second2.
Despite recent helmet improvements (mostly better cushioning in the liner area to better reduce head acceleration levels), concussions seem to continue unabated, so the various helmet improvements have not significantly helped to reduce the number of occurrences. One likely reason for the lack of success in reducing concussions is that the helmet improvements made so far have mostly concentrated on the linear acceleration issue, and have mostly or completely ignored the angular acceleration issue.
The lack of real reductions in concussions may be the result of a simple misconception about what goes on inside the head to cause a concussion. The simplified view is that when the skull is stopped too abruptly, in say a frontal impact, the brain continues on to strike the inside of the skull at the front, and if the impact is severe enough the brain can even rebound and strike the inside of the skull at the rear. The former is termed a coup injury and the latter a contrecoup injury. As a result of the above simple explanation, the main object in making helmet improvements has been to stop the skull less abruptly, i.e., taking steps to reduce its linear deceleration. That is what most of the recent helmet improvements have concentrated on doing. Yet it will be herein shown that nature's own thin layer of cerebrospinal fluid or CSF between the brain and the inside of the skull is extremely effective through its buoyancy effect in mitigating the envisioned impact between the brain and the front of the skull in an abrupt linear stop, even at head deceleration levels that greatly exceed 79 Gs. So, contrary to current thinking, high linear acceleration, or deceleration, does not provide the entire picture, and one needs to look further, particularly at the angular acceleration of the head.
But angular acceleration is not part of that simplified picture of what happens to the brain in a concussion, so it tends to get ignored. And yet, unlike with linear acceleration, the cerebrospinal fluid is not as effective in eliminating damaging internal impacts of the brain against the inside of the skull in response to an abrupt high angular acceleration of the head. Two contributors to angular acceleration are herein identified which may either add or subtract depending on the direction of the impact and its location, both with respect to the neck position as will be discussed below. Limiting the linear acceleration or deceleration of the head, which current helmet designs do fairly effectively, is helpful in limiting the first contributor to angular acceleration, which is the pendulum motion of the head and neck together. But the current helmet designs do little or nothing to limit the second contributor to angular acceleration, which is the rotational motion of the head at the top of the neck. If this second contributor to angular acceleration could be limited as well, it would go a long way toward reducing the high levels of angular acceleration that appear to lead to concussions. Indeed, the field data show that without this second contributor to angular acceleration, most of the current concussion level football impacts would fall short of the accepted threshold concussion level for angular acceleration. Accordingly, the overall number of football concussions may be significantly reduced if a new helmet design that could additionally significantly lower this second angular acceleration contributor were to be widely implemented.
Note that regarding the terminology used in the preceding and following discussion and throughout the specification, on occasion the terms acceleration and deceleration are used within their specific intended meanings, but usually the two terms may be interchanged, so when the term acceleration is used it applies equally well to a deceleration and vice versa. Also, within the specification, the terms angular, rotational, circumferential, tangential, and lateral are often used interchangeably, as are the terms linear, radial, centered, straight-on, and normal. The term off-center refers to any direction between centered and tangential. Finally, the terms radial and radially should be interpreted as meaning substantially radial, as it usually relates to a non-spherical surface (object) such as a spheroid, ellipsoid, or ovoid surface.
To understand how the present invention addresses the concussion problem, it is helpful to first review the results of some comprehensive in-situ football data. In a study conducted by Virginia Tech in 2007, and reported on by Rowson, et al, in the Journal of Biomechanical Engineering, June 2009, Vol. 131, ten six-degree-of-freedom (6DOF) instrumented helmets were used to collect data during both practices and games on offensive and defensive linemen. These biggest players wear the largest helmets which are able to accommodate the instrumentation. Each 6DOF system consists of 6 dual axis micro-electro-mechanical-system (MEMS) accelerometers for a total of 12 independent outputs (a minimum of 9 are needed in a 3,2,2,2 configuration so the extra 3 outputs provide for some redundancy) installed in a Riddell Revolution model football helmet (a recent design for concussion avoidance), a wireless transceiver, and an on-board memory for up to 120 impacts with 8 bit resolution data being acquired continuously at a sample rate of 1,000 Hz per channel. A data set was triggered and saved when any accelerometer experienced an impact level of 10 Gs or more. Impact data sets are 40 milliseconds long (8 ms pre-trigger and 32 ms post-trigger). All of the saved data was transmitted to the sidelines by a commercial computerized helmet impact transmission system, called HITS, to be further analyzed. All of the MEMS miniature accelerometers were held tightly against the skull of the helmet wearer by the foam padding of the helmet to help insure good skull motion data, and the raw data was combined in the following coordinate system: The positive x-axis is directed out of the face (perpendicular to the coronal plane), the positive y-axis is directed out of the right ear (perpendicular to the midsagittal plane), and the positive z-axis is directed out of the bottom of the head (perpendicular to the transverse plane). The origin approximates the center of gravity (c.g.) of the head.
In all, 1712 impacts were recorded, 570 during games, 1142 during practices. Although 11 peak linear accelerations exceeded 80 g and 12 peak angular accelerations exceeded 6,000 rad/sec2, no instrumented player sustained a concussion during the 2007 season. The maximum recorded peak linear acceleration was 135 g and the maximum recorded peak angular acceleration was 9,222 rad/sec2, each over 50% more than accepted NFL threshold values. However, in other studies, players who experienced lower values than the NFL threshold values did sustain concussions. Clearly, the situation is far more complex than just the levels of peak acceleration.
FIG. 1 shows an average linear acceleration response in the Virginia Tech in-situ data. The average peak acceleration value was 23 g and all the acceleration/deceleration waveforms lasted approximately 14 milliseconds as shown. For the larger accelerations (and the larger angular accelerations), the timing remained approximately the same.
FIG. 2 shows a scatter plot of the change in linear velocity of the head vs. peak linear acceleration for all of the impacts. Only a few impacts represented a change in velocity of up to 20 ft/sec and the vast majority of the rest were less than half that value. Despite a slight offset about the origin, note the approximate linear relationship between change in velocity and peak linear acceleration.
FIG. 3 shows a scatter plot of the change in angular velocity of the head vs. peak angular acceleration for all of the impacts. Again note the approximate linear relationship.
FIG. 4 shows a scatter plot of peak angular acceleration vs. peak linear acceleration for all of the impacts. Note that each impact results in both a linear and an angular acceleration. The reference line is 4,300 rad/sec2 per 100 Gs. But there is little evidence of linearity or correlation between the two accelerations. That is, there can be high angular acceleration at the same time as low linear acceleration, and vice versa. How this can physically happen provides the clue for how to keep the peak angular acceleration value below the concussion threshold value in most cases. As will be discussed, the peak angular acceleration value is what is most damaging to the brain, but the peak linear acceleration value, although not particularly damaging in its own right, is still very important in its role as a contributor to the peak angular acceleration. This apparent dichotomy with respect to the role of peak linear acceleration has likely led to the confusion that's existed among current researchers trying to determine the significance of peak linear and angular accelerations in concussions.
Before attempting to fully understand FIG. 4, we need to first explore the head, neck, and body connection. In all head impact cases the forces and torques that eventually halt the impulsive and inertial motions of the head must arise from the more massive body and these forces and torques come through the neck. If the neck were so rigid that the head could not move at all with respect to the massive body, it would be unlikely that any football player could receive enough linear or angular acceleration to cause a concussion. Thus one can assume the stronger the neck connection to that massive body (the stronger the neck muscles), the lesser the impulsive inertial motions of the head will be. That may be why professional football players, who have stronger necks than high school players, do not suffer proportionally more concussions even though they are hit harder. Also, the striking (hitting) players in a collision appear to suffer fewer concussions than the struck (hit) players and one reason might be because the striking players may have tensed their neck muscles in preparation for the impact while the struck players may be caught unawares. Another reason is presented later when it can be better understood.
But since no football player's neck is totally rigid, the allowed motions need to be considered to better understand FIG. 4, with its non-correlating angular and linear acceleration levels. The neck contains seven cervical vertebrae that connect the skull to the thoracic vertebrae and the rest of the body. The neck can curve one way at the top by the head and another way at the bottom where it joins the more massive body. At the bottom, the neck can bend forward toward the chest or backward toward the back, and also it can bend toward the right shoulder or toward the left shoulder. At the top of the neck (pivoting at about ear level as viewed from the side), the head may independently rotate in any of three planes: first, the shaking of one's head in a vertical midsagittal plane “yes” motion; second, the shaking of one's head in a horizontal transverse plane “no” motion; and third, the cocking of one's head left or right in a vertical coronal plane. As will be shown below, the independent rotation of the head at the top of the neck is the main reason for seeing wildly different angular and linear accelerations in a given impact.
Based on the above-described allowed head-neck motions, in order to analyze what is going on it is useful to envision the head-neck system as an “apple-on-a-stick,” where the stick (the neck) is able to pivot in two directions (forward and backward and side to side) at its base (where it joins the body) thereby enabling a sort of pendulum motion, and the apple (the head) is able to pivot in all three directions at the top of the stick (in other words: at the top of the neck, at about ear height) thereby enabling an additional rotational motion of just the head. The first motion (the head-neck pendulum motion) contributes to both the linear and the angular acceleration of the head, while the second motion (the rotational motion of just the head at the top-of-the-neck) contributes mostly to just the angular acceleration of the head. These two contributors to angular acceleration, when existing in the same plane, may either add or subtract depending on the direction of the impact and its location, as will be discussed below. When in different planes, the two contributors to the total head angular acceleration also combine but not in a direct fashion. Limiting the linear acceleration or deceleration of the head in response to an impact, which current helmet designs do fairly effectively, is helpful in also limiting the first contributor to head angular acceleration, the head-neck pendulum motion. But current helmet designs do very little to limit the second contributor to head angular acceleration, the independent top-of-the-neck rotational motion of the head. That fact is evidenced by how easily a player's head can be jerked around, for example, when another player yanks his facemask.
It is a fundamental assertion of the present invention that high angular acceleration of the head is the primary causer of brain injury in a head impact, and, conversely that high linear acceleration of the head is not the main injury causer, except through its contribution to head angular acceleration via the previously described head-neck pendulum motion. At the heart of this assertion largely vindicating linear acceleration is the contention that, contrary to popular belief, when the skull is suddenly stopped in a helmet-to-helmet collision, the brain does not continue on unimpeded to crash against the inside of the skull in the direction of the impact, then to potentially rebound to crash against the inside of the skull in the opposite direction as well. Moreover, this contention is a fact, as will be shown in the following paragraphs.
It was previously stated, without supporting evidence, that the buoyancy of the brain in the surrounding cerebrospinal fluid is very effective in eliminating an impact of the brain against the inside of the skull wall (the cranium) in very high linear acceleration and deceleration (impact) situations. The following examples and discussion provide the supporting evidence to confirm the foregoing statement.
Picture a car crashing head-on into a concrete wall. The car's inhabitants (assuming no seat belts and no air bags) will continue to move forward until they smash into one or more of the inside structures of the car (dashboard, windshield, etc.) That is how a concussion is typically described, where the skull plays the role of the car and the brain plays the role of its inhabitants. However, what if the car were filled with water instead of air, and the inhabitants (now properly fitted with SCUBA gear) are neutrally buoyant in the water, like the brain is approximately neutrally buoyant in the surrounding cerebrospinal fluid. Now upon the collision of the car into the immovable wall, the car, the water, and the inhabitants all come to a stop in short order and none of the inhabitants smash into the windshield or other interior car surfaces. Why?
By the well proven Equivalence Principle in physics, inside a small windowless room in outer space nothing can tell the difference between an acceleration/deceleration force and a gravity force. Thus, if the deceleration of the car were a constant 1 G, that would be equivalent to simply standing the car on end, front side down, on Earth. In that case, all of the inhabitants in the water-filled car would remain as neutrally buoyant as they were before, suspended in-place like a submarine in the ocean, and no one would crash downward into the windshield or other interior surfaces of the car. If the deceleration were a constant 100 Gs, that would be equivalent to standing the car on end on a planet with 100 times the gravity of Earth, and again everyone would remain neutrally buoyant, suspended in-place, and no one would crash into the windshield. Physically, a linear pressure gradient is formed in the water. On the 1 G Earth, in every body of water, no matter how big or how small, the linear pressure gradient goes from zero at the top surface (plus atmospheric pressure) to a pressure at the bottom equal to the weight density of the water (its mass density times the acceleration of gravity) times the depth of the water (plus atmospheric pressure). For a neutrally buoyant object in the water, the effective pressure gradient (along the object) times the effective area of the object (acted on by the pressure gradient) exactly counters its weight (its mass times the acceleration of gravity). At 100 Gs, the weight of the object is 100 times as much, but the weight density of water is also 100 times as much so the effective pressure gradient is 100 times as much and the object remains neutrally buoyant, and stationary. This is equivalent to what happens under acceleration.
It is not necessary to just accept this at face value. It can be verified experimentally using a 1 inch diameter solid polystyrene ball which has a specific gravity of 1.040, and a 5.5% saline solution of water which has a specific gravity of 1.040 at 68° F. Place the ball and saline solution in a 2 inch diameter transparent hard plastic tube closed and sealed at both ends. Make sure all the air bubbles have been removed. Then with the ball suspended in the middle of the tube, smack the tube axially into a hard stationary surface as hard as possible and observe how the ball moves. See if the ball which represents a neutrally buoyant brain, suspended in the saline solution which represents the cerebrospinal fluid, crashes into the front impact surface of the tube representing the inside of the skull. It should not. Indeed if what has been stated above is correct—and it is—the neutrally buoyant ball should not move at all—and it doesn't.
When talking about the brain, however, the brain is not exactly neutrally buoyant in the surrounding cerebrospinal fluid. It is about 3% more dense than the fluid. So the brain will continue to move forward when the forward-moving skull is abruptly decelerated to a stop, but by how much and with what remaining velocity?
Picture a non-helmeted man running through a darkened space with his head held well forward when suddenly his head strikes a wall while he's running at, for example, 10 ft/sec (which is about an 8 minute mile pace). The key constraint in this example is that the orientation of the man's skull remains unchanged throughout the process, so that there is no angular acceleration. Also, it is assumed the man is fortunate enough to not break his neck, nor fracture his skull, but his skull's limited elasticity when combined with the stiffness of the wall will stop his skull in (say) just 2 milliseconds (a reasonable assumption). We can further simplify the analysis by assuming, in addition, that the deceleration of his skull is constant over those 2 milliseconds, and with that assumption the resulting calculated deceleration will be 155.3 Gs. Note that the peak deceleration would be higher without that assumption.
Now what happens to the man's brain at the same time? His brain weighs about 3.1 lbs and approximates a 6.8 inch long top-half semi-ellipsoid or ovoid. The weight density of his brain is about 0.0375 lbs/in3, and the weight density of the cerebrospinal fluid which surrounds it is about 0.0364 lbs/in3. The cerebrospinal fluid CSF decelerates along with the skull resulting in a linear pressure gradient in the CSF (for those 2 milliseconds) that ranges from zero psi gauge pressure at the back of the brain to 38.4 psi gauge pressure at the front of the brain where the skull was impacted (6.8×0.364×155.3=38.4). Thus, acting upon each small segmental surface area of the brain, there is a front/back force on that brain area segment equal to the front/back projection of the area segment times the gauge pressure at that location. This calculation yields a resultant decelerating force of 466.5 lbs. with the resulting deceleration of the 3.1 lb brain being 150.5 Gs. Thus the brain is significantly slowed along with the skull, but not quite as much as the skull.
The distance the man's skull travels during the deceleration is:dsk=V0t−½askt2  (Equation 1)where V0=10 ft/sec; t=2 msec; ask=155.3 Gs→dsk=0.120 inches
The distance his brain travels during the deceleration is:dbr=V0t−½abrt2  (Equation 2)where V0=10 ft/sec; t=2 msec; abr=150.5 Gs→dbr=0.124 inches
Thus during those 2 milliseconds of deceleration, the man's brain closes the gap between itself and the front of his skull by only 0.004 inches (about the thickness of a piece of paper). The initial gap is about 0.100 inches (approximately 2.5 mm), consisting of the outer hard dura mater layer, the inner soft pia mater layer which covers the brain, and the filament-like arachnoid layer and the CSF-filled subarachnoid space in between.
So, at the end of the 2 millisecond skull deceleration period, the speed of the man's skull is 0 ft/sec and the speed of his brain is all the way down to 0.31 ft/sec (from 10 ft/sec). In terms of energy, due to kinetic energy's speed squared relationship, 99.9% of his brain's initial kinetic energy has already been dissipated, leaving just 0.1% of its initial kinetic energy to yet be dissipated. Since the cerebrospinal fluid is no longer decelerating to provide a decelerating force through an acceleration induced linear pressure gradient, the deceleration must be accomplished by squeezing more of the cerebrospinal fluid out of the remaining 0.096 inch space and compressing the compressible pia mater and arachnoid layer. The remaining required deceleration of 0.19 Gs, which corresponds to a decelerating force of only 9.3 ounces, is not very likely to be damaging.
Before knowing the above analysis one would have assumed that a 155 G deceleration impact on the skull would certainly cause a concussion. In light of the above analysis, however, that seems to no longer be the case, even for a head deceleration level more than two times what the NFL considers to be the linear acceleration/deceleration threshold level for concussions (79 Gs). Why then does a high peak linear acceleration level of the head matter? (Recall that in the above example, the orientation of the cranium was held constant, so there was no angular acceleration of the head.)
For real-life impacts, however, high linear acceleration levels usually do matter because through the previously described head-neck pendulum motion, the linear acceleration of the head also contributes to the angular acceleration of the head. When the linear acceleration component perpendicular to the neck at the c.g. of the head (located about 8 inches from the lower neck pivot) is at a level of 79 Gs, its contribution to the resulting angular acceleration of the head is 3,816 rad/sec2. That is just two-thirds of the NFL threshold angular acceleration level of 5,757 rad/sec2. Moreover, only rarely will a measured 79 G peak linear acceleration level occur in a direction perpendicular to the neck (at the c.g. of the head), so in order to attain a 79 G perpendicular component the total peak linear acceleration level would normally need to be even higher. But in order to reach the angular acceleration concussion level, there will usually need to be not just a high peak linear acceleration level (to yield a reasonably high angular acceleration value through the head-neck pendulum effect), there needs to also be a significant and additive head rotational acceleration component present as well. This is the previously mentioned top-of-the-neck second head rotational acceleration component—the one the present invention attempts to further reduce.
To reinforce all the above and put the numbers in prospective, a second football study is presented. This study, reported on by Broglio, et al, in Medicine and Science in Sports and Exercise, 2010, followed 78 high school football players wearing Riddell Revolution helmets instrumented with the previously described Head Impact Telemetry System, (HITS) through four seasons of practices and games from 2005 to 2008. In all, 54,247 impacts were recorded (the impacts triggered whenever one of the accelerometer channels from the six dual axis units exceeded a threshold of 15 Gs). The data included 13 impacts that resulted in concussions. The recorded average peak linear acceleration levels were about 26 Gs, and the average peak angular acceleration levels were about 1,600 rad/sec2, very similar to the previously cited data. But this study is more valuable because it includes data from actual concussion-causing impacts. From the data, the authors developed a concussion predictor “tree.” The tree starts off not surprisingly with an angular acceleration threshold question.
1st Question: Angular Acceleration>5,582 rad/sec2                 Answers: (No—53,563 impacts, 0 concussions)—0%        (Yes—684 impacts, 13 concussions)—1.9%        
↓(yes)
2nd Question: Linear Acceleration>96 Gs                Answers: (No—525 impacts, 2 concussions)—0.4%        (Yes—159 impacts, 11 concussions)—6.9%        
↓(yes)
3rd Question: Impact location; front, side, top                Answers: (No—77 impacts, 0 concussions)—0%        (Yes—82 impacts, 11 concussions)—13.4%        
↓(yes)
4th Question: Angular Acceleration<8,845 rad/sec2                 Answers: (No—35 impacts, 1 concussion)—2.9%        (Yes—47 impacts, 10 concussions)—21.3%        
↓(yes)
5th Question: Linear Acceleration<102 Gs                Answers: (No—38 impacts, 5 concussions)—13.2%        (Yes—9 impacts, 5 concussions)—55.6%        
For the 13 concussion causing impacts, the key metric was the resultant peak angular acceleration level. A minimum level of 5,582 rad/sec2 was the indicated value, but the mean level was 7,229 rad/sec2. The indicated minimum level of angular acceleration was a necessary, but not sufficient condition for the 13 concussive impacts (out of 54,247 impacts). From the standpoint of identifying better helmet protection, identifying a necessary condition is paramount, but from the standpoint of identifying a predictive metric, the necessary condition is not enough. In other words, 98% of the time (671 times out of 684 times), a player who received an angular acceleration greater than 5,582 rad/sec2 did not suffer a concussion. So angular acceleration is a poor predictor. However, no player suffered a concussion as a result of receiving any of the 53,563 impacts where the angular acceleration level was less than 5,582 rad/sec2. That is a powerful protection identifier—i.e., to simply incorporate a protective measure that will keep the head angular acceleration level below 5,582 rad/sec2 as much as possible.
A key point previously made, now bears repeating. For those special cases that exhibit no local rotation of the head at the top of the neck, (envisioning all the motion of the head as just a pendulous apple on a stick pivoting at the base of the neck), a linear acceleration of the head still results in an angular acceleration of the head. For a=79 G, and r=8 inches, angular acceleration α=3,816 rad/sec2. So for this very simplified case, a supposed concussion level for linear acceleration does not result in a concussion level for angular acceleration. To reach the concussion level for angular acceleration, there must also be a local angular acceleration (one that causes a local rotation of the head at the top of the neck) that adds to the above pendulum angular acceleration and the total combined angular acceleration value is the true culprit. The fact that in the first study's data (the college study), the measured angular accelerations were all over the map as compared to the measured linear accelerations (FIG. 4) is proof that local rotational accelerations of the head of the same order of magnitude as the head-neck pendulum head angular accelerations exist, and may occasionally fully add or fully subtract from the latter. From the above numbers, without the local angular acceleration contributor (to rotate the head at the top of the neck) it would take a pure 120 G linear acceleration to result in a pendulum angular acceleration that exceeds the 5,757 rad/sec2 NFL threshold concussion value. Thus it should be clear that if the local rotational angular acceleration contributor could be eliminated (or significantly reduced) by the design of the helmet, then the pendulum angular acceleration all by itself would rarely be able to cause a concussion in a helmeted football player.
All of the concussed high school football players in the study not only received high resultant peak angular acceleration levels but also high resultant peak linear acceleration levels (the lowest was 74 Gs). But apparently many received the latter without the former and did not get concussions. The mean resultant peak linear acceleration level for the concussed players was 105 Gs. Assuming an average angle of 45° with the neck for the impact direction, and with the cosine of 45°=0.707, that would yield an average component perpendicular to the neck axis of 74 Gs, which by the previously described head-neck pendulum motion would yield a corresponding peak angular acceleration level of 3,575 rad/sec2. That is approximately half the indicated mean level of 7,229 rad/sec2 which the concussed players received, so on average, only about half the resultant peak angular acceleration for those 13 concussed players is the result of the linear acceleration acting through the head-neck pendulum motion. The other half—at least another 3,600 rad/sec2 on average—must have come from the purely rotational acceleration of the head at the top of the neck that the present invention is intended to reduce.
A head angular acceleration threshold has been identified below which players seem not to get a concussion. Yet above that threshold they get a concussion only 2% of the time. Why? Does the cerebral spinal fluid CSF still play some sort of protective role for angular acceleration as it does for linear acceleration?
It was previously shown how the brain's near-buoyancy in the CSF causes a rapid pressure gradient rise in the CSF in synch with and proportional to the skull's rise in linear acceleration/deceleration, with the maximum pressure occurring at the impact location, and it was also shown that the pressure gradient increase causes an almost matching acceleration/deceleration of the brain, so no significant impact of the brain occurs against the inside of the skull. Indeed, researchers using tiny pressure transducers implanted in the brains of cadavers for head impact tests have recorded pressure waveforms near the impact location that exactly match the linear acceleration waveforms of the decelerating skull. Some researchers, who did not appreciate the fact that what they were recording was the brain's protective mechanism against linear acceleration, have conjectured that perhaps the rapid pressure increase is the damaging mechanism. But studies have shown that the brain is not damaged by compression, only by stretching, shearing, or twisting. Since the brain is not being bounced back and forth as commonly pictured, it must be the sudden rotation of the head that is causing the cranium (the portion of the skull that surrounds the brain) to impact the brain at one or more locations which results in that stretching or twisting. However, because the cranium and the brain are not spherical, but instead semi-ovoid and oblong, at the oblong extremities an angular acceleration can resemble a transverse linear acceleration and as a result the CSF can experience quasi-linear acceleration induced pressure gradients at the oblong extremities which tend to gently (over a wide surface area) rotate the near neutrally buoyant brain along with the cranium, and so the CSF is still partially protective against angular acceleration induced internal impacts, just not nearly as effectively as for pure linear accelerations. Just how protective this will be can depend on a host of factors including but not limited to: the cranium and brain's different oblong nature in the different axes, individual physical shape differences, how the brain's undulating surface high regions and low regions line up with the major angular acceleration axis, and how variations in the thickness of the CSF layer locally line up at potential rotational impact points. With all that variability, it is perhaps not surprising that 6,000 rad/sec2 might result in a concussion in one instance, but 9,000 rad/sec2 might not result in a concussion in another. It is also not surprising that the CSF would be partially protective against head angular acceleration; otherwise we might all be giving ourselves concussions every time we shake our heads yes or no.
In a concussion the cranium pushes on the surface of the brain at just a few points which then bear the brunt of having to push the entire jello-like brain mass around to try to follow the sudden cranial motion, and so these points experience the most localized strain and shearing and may suffer the previously cited coup and contrecoup injuries. Thus the coup and contrecoup injuries should not be visualized as a one-two punch caused by the brain first crashing against the inside of the cranium at the “front” then rebounding to later crash at the “rear,” but rather as a virtually simultaneous, locally stressful and strain-full pushing of the brain around at a few widely separated points where it comes into contact with the cranium. And when a concussion occurs these are not as much physical injuries as they are chemical events wherein the momentary stretching of the walls of the brain cells enables potassium ions to suddenly escape and be replaced by calcium ions, which is a very negative event that may take days or even weeks to correct itself. While being pushed around rotationally, the internal regions of the brain may also get stretched and sheared, which, and as noted above, more than any simple compression is what most agree causes serious brain injury. The most severe form of injury is called Diffuse Axonal Injury, or DAI. DAI damage occurs mostly at the juncture between the outer grey matter and the slightly more dense inner white matter toward the brain's interior, as any angular relative motion between the two could stretch and tear the interconnecting axons over a wide ranging (highly diffuse) area. Some brain experts say that at least some degree of DAI is present with any concussion that involves a loss of consciousness. Strain levels (and high strain rates) of more than 10% are considered to be almost always damaging. Indeed the highest degree of correlation to concussion seems to be the product of brain tissue strain and strain rate, something nearly impossible to measure on football players in situ. But from the standpoint of inventing a more protective helmet (against concussions), it is not necessary to understand all the possible damaging or mitigating factors that exist when translating a high peak angular acceleration level into a high product of strain and strain rate in the exterior and interior regions of the brain.
The liners of most current football helmets already effectively reduce the linear acceleration of the head as compared to the linear acceleration of the helmet shell, which in turn reduces any head angular acceleration contribution that arises through the head-neck pendulum effect. But current helmet liners are not designed to reduce the rotational acceleration of the head that arises from the rotational acceleration of the helmet shell, and this rotational acceleration (from both of the above discussed studies) contributes directly to the total angular acceleration level of the head. Thus, one way to create a better concussion-reducing helmet is to make the helmet liner also reduce any rotational contributor to the total peak angular acceleration of the head which are coming from the rotational acceleration of the helmet shell. Note that for helmet impacts, it is far more likely for a wearer to experience a sudden angular acceleration than an angular deceleration, although the same result would occur either way.
Looking at the shiny, round, hard plastic surface of a football helmet it may be hard to imagine how a helmet shell can even acquire a large rotational acceleration in a helmet-to-helmet collision. After all, it is so smooth and has a rounded, low friction surface. If one holds two empty football helmets by their facemasks, and bangs them together, they just bounce away with little resulting rotation. So one's initial conclusion may be to assume that all of the forces always lie along a line of contact normal to the two surfaces at their contact point, and thus aren't able to cause any rotation. But that is just for the special case where the initial relative motion also lies along the line of contact. If one bangs the helmets together off-center (not along their line of contact), a totally different story emerges—there is a lot of rotation, even without much friction between the two smooth surfaces. The reason is there is still a normal force component that dimples each helmet shell inward (very significantly) at the point of contact. What is amazing is how rapidly the diameter of the dimpled-in area (an effective flat from the standpoint of the other helmet) can increase, and thereby have its effect brought into play. And its effect, in conjunction with any relative tangential velocity, is to cause a suddenly increasing rotation of each helmet shell with accompanying high rotational acceleration levels.
Take the case of two football players running or diving at each other at a closing speed of 25.6 ft/sec (or 7.8 msec which is faster than any in the college study, see FIG. 2), and then impacting helmet-to-helmet, not in a centered collision but in a 45 degree off-center collision. Therefore, their effective relative speed in both the helmet normal direction and the helmet tangential direction is about 18 ft/sec (cos 45°=0.707). Assume the helmet shells' normal speeds are shared 50/50 at 9 ft/sec each and each's normal motion is stopped in 5 milliseconds with an assumed approximate quarter sine wave decelerating force. The calculated resulting normal displacement of each helmet shell (equal to the dimpling-in distance) is approximately 0.3 inches, which corresponds to an elastically flattened diameter of 3.2 inches (a little wider than a hockey puck). In this example, the elastic flattening that takes place in 5 milliseconds returns to its original shape in another 5 milliseconds, after which the shells lose contact and separate. Note that in the normal direction both the helmet shells and the players heads are accelerated/decelerated for the full 10 milliseconds that the helmet shells remain in contact. It can be assumed with no loss in generality that the shells came together with equal speeds then decelerated to zero speed in 5 milliseconds, and then in the next 5 milliseconds they were accelerated back up to separation speeds equivalent to their speeds at initial contact but in the opposite directions. Meanwhile, thanks to the liners, the heads may take advantage of the full 10 milliseconds to decelerate to a stop and then the heads (via the neck muscles) can decelerate the shells back to zero speed at lower acceleration levels over a longer time after the shells lose contact with each other. To the heads, that looks like a continued low level acceleration in the same direction as during contact, which is the reason for the long descending plateau region of FIG. 1.
Events occurring within ten milliseconds may be too fast to be seen by the human eye. However, that is not too fast for some of the 18 ft/sec differential tangential velocity in the above non-centric impact example to be picked up by both helmet shells. They'd be tangentially accelerated in the same rotational direction by an oppositely directed friction force exerted on each by the other which is generally proportional to their shared oppositely directed normal force, so the resulting angular acceleration might be expected to have the same sort of waveform as the linear acceleration and be synched to it. If the two shells share that tangential velocity gain equally, then each 9 inch diameter helmet shell could pick up a circumferential velocity of up to 9 ft/sec, which using the same waveform characteristic and same timing would correspond to a maximum peak top-of-the-neck angular acceleration component of up to about 4,000 rad/sec2. That value is right in the ballpark of what might be expected to encompass the actual value for an off-center impact of that intensity, and is consistent with most of the cited football data. The resulting calculated circumferential displacement of the helmet shell is less than half an inch. That establishes the design parameter for what must be accommodated in terms of relative circumferential displacement between the outer shell and the head cap (i.e., by the liner) at not more than an inch.
Note, for those impacts that are near a full 90 degrees off-center (a grazing impact) the relative tangential speed component may be very high, but the normal speed and force components are very low by comparison, so the dimpling-in is small and the time to take-on the tangential speed (via any tangential force) is also small. Also for impacts that are near 0 degrees off-center (a near normal impact) the normal speed and force components may be very high and the dimpled-in time may be also high, but the relative tangential speed is very low by comparison so the tangential speed that can be taken on is limited.
The present invention provides an improved helmet system which contains three essential parts: an inner head cap that is attachable and detachable to the head of a user and moves with the head; an outer impact resistant hard shell which moves independently from the head cap and user's head; and a returnable, energy absorbing liner located in-between the head cap and the outer shell which is compliant both radially and circumferentially in all directions. The returnability feature may be manual for use in sports or other activities where the expected impacts are rare such as bicycling, but automatic for use in sports or other activities, such as football, where the impacts are numerous and repetitive.
The preferred embodiments of the present invention employ an energy absorbing viscoelastic polymeric foam material (PU, EVA, EPP, or the like) to form the liner between the outer shell and the head cap. The liner is configured to be able to reduce linear accelerations and decelerations of the head compared to those of the outer shell as effectively as current prior art helmets. In addition, with the present invention the viscoelastic polymeric foam material of the liner is specially configured to be able to reduce angular accelerations of the head compared to those of the outer shell. To not compromise the latter function, the chin strap with its attached chin protector is fastened to the head cap, which is conformal to and moves with the head, and the chin strap is not fastened or otherwise attached to the outer shell, which has been enabled by the special configuration of the connecting viscoelastic polymeric foam material to be able to move relative to the head cap and the head both linearly and angularly. After an impact, where the outer shell has moved linearly and angularly relative to the head cap and the head, the specially configured liner either causes the outer shell to automatically return to its initial pre-impact start position relative to the head cap and the head, or it enables that return to be manually completed.
In a first preferred embodiment of the present invention, wherein the return is automatic, the special configuration of the viscoelastic foam liner is comprised of a plurality of side-by-side, long and narrow foam columns with their long sides generally radially-oriented so they are slightly tapered (with their wider ends outward). The long narrow foam columns span and nearly fill the space between the outer surface of the head cap and the inner surface of the outer helmet shell, with each column being adhered at each end to each surface. The cross sections of the columns may be triangular, rectangular, pentagonal, hexagonal, round, oval, or other suitable shape, but in all cases should have sufficiently effective length-to-width ratios for the necessary transverse compliance, in addition to the necessary linear compliance, which gives the liner the ability to reduce the angular accelerations of the head.